Cardiovascular disease remains the leading cause of death in the United States, accounting for nearly 1 million deaths in 1996. Of these fatalities, 50% are attributed to coronary artery disease that arises from low-density lipoprotein (LDL) cholesterol, which transport about 75% of the cholesterol. It can penetrate the artery wall where it interacts with free radicals that attack and modify its form. The resulting oxidized form of LDL triggers white blood cells in the immune system to gather at the site, forming thick substance called plaque and causing inflammation. The plaque will build up eventually constricting the walls, in the process known as atherosclerosis [1].
The second most common heart operation in the western world is heart valve replacement [2]. The main types of replacement valves for heart valve replacements are mechanical and bioprosthetic, both with advantages and disadvantages. Mechanical heart valves are made of non-biologic materials, and their advantages are their durability and structural reliability. Their main disadvantages are the patient risk of thromboembolism due to the poor blood compatibility and flow abnormalities. To reduce the risk, the patient requires lifetime anticoagulant therapy [3, 4]. Bioprosthetic heart valves are made in part of animal tissue, thus maintaining a low level of thromboembolism without the need of long-term anticoagulant therapy. They also have improved hemodynamics because their flow pattern is similar to natural valves. However, their major disadvantage is their limited durability, due to structural dysfunction from calcification and noncalcific tissue deterioration. More than 50% of them fail between 10 to 15 years and require re-operation [3, 4, 5].
One of the most common treatments for coronary artery disease is coronary artery bypass surgery, which is the revascularization of the damaged myocardium [6]. Normally, a suitable length of the patient's saphenous vein is used to provide blood to the heart tissue. The main disadvantage is “vein graft disease”, which is the deterioration and occlusion of the vein graft due to further advancement of the patient's coronary artery disease [7, 1].
Therefore, here lies the need to develop a material that will not only display similar mechanical properties as the tissue it is replacing, but also shows improved life span. One promising class of materials are hydrogels.
Hydrogels
Hydrogels are hydrophilic polymer networks produced from reactions of one or more monomers or by association bonds between chains that can absorb from at least 20% to up to thousands of times their dry weight in water [8, 9]. Hydrogels may be chemically stable or they may disintegrate and dissolve with time. They are called either physical (reversible) or chemical (permanent) hydrogels. Physical hydrogels have networks held together by molecular entanglements and/or secondary forces such as hydrogen bonding, van der Waals interactions, ionic or hydrophobic forces. Physical hydrogels are not homogeneous due to regions of high crosslinking density and low water swelling, called clusters, dispersed within low crosslinking density and high water swelling, or hydrophobic or ionic domains that create inhomogeneities. Chemical hydrogels are covalently crosslinked networks, but they may also be generated by crosslinking of water-soluble polymers, or by converting hydrophobic polymers to hydrophilic polymers. Chemical hydrogels are also not homogeneous due to clusters of molecular entanglements. Chain loops and free chain ends also produce network defects in both physical and chemical hydrogels, and they do not contribute to the permanent network elasticity [8, 10].
An important characteristic of hydrogels is their swelling behaviour in water, since after preparation they have to be in contact with water to yield the final solvated network structure. Highly swollen hydrogels are those of poly(vinyl alcohol) (PVA), poly(ethylene glycol), and poly(N-vinyl 2-pyrrolidone), among others. Poly(vinyl alcohol) (PVA) is a hydrophilic polymer with various characteristics desired for biomedical applications, such as high degree of swelling, uncomplicated chemical structure, rubbery/elastic nature, and non-toxic. PVA can be converted into a solid hydrogel by crosslinking. Crosslinking can be accomplished by using several methods. For biomedical applications, physical crosslinking has the advantages of not leaving residual amounts of the toxic crosslinking agent, and higher mechanical strength than the PVA gels crosslinked by either chemical or irradiative techniques. The mechanical properties of the PVA hydrogels are similar to that of soft tissue, including elasticity and strength, and can be controlled by changing the number of thermal cycles, PVA concentration, thawing rate of the thermal cycling process, and freezing holding time among other parameters [11, 12, 13]. A PVA based bioprosthetic heart valve stent has been fabricated. However, the mechanical strength and stiffness of these PVA materials were weak and did not fully match the mechanical properties displayed by the cardiovascular tissues such as arteries and heart valves.
Poorly swollen hydrogels are those of poly(hydroxyethyl methacrylate) (PHEMA) and its derivatives. However, the desired swelling properties can be achieved by copolymerization of a hydrophilic monomer with a less hydrophilic one. This gives a vast range of swellable hydrogels, and the swelling characteristics are of great importance for biomedical and pharmaceutical applications. This equilibrium degree of swelling affects the solute diffusion coefficient through these gels (control release applications), the surface properties and mobility (coating applications), the optical properties (contact lenses applications), and the mechanical properties of the hydrogel (tissue replacement applications) [14].
The main areas in which hydrogels are used as biomaterials is in contact lenses, synthetic wound coverings, drug delivery systems, organ and tissue replacements, and permselective membranes [8, 14, 10, 15, 16, 11, 17, 18, 19, 5, 20, 13]. One of the major disadvantages of hydrogels is that when dehydrated, they are hard and brittle, but when swollen in water, they become rubbery with a very low tear and tensile strength. This has a profound effect on the life span of the lenses. Most of the research tries to improve the mechanical properties by looking at a variety of polymer combinations and cross-linking agents, such as acrylamide and acrylonitrile-based hydrogels, and vinyl pyrrolidone copolymers [21].
PVA has a relatively simple chemical formula with a pendant hydroxyl group and a crystalline nature, which allows it to form a solid hydrogel by the crosslinking of the PVA polymer chains. Vinyl alcohol (monomer) does not exist in a stable form and rearranges to its tautomer, acetaldehyde. PVA is produced by free radical polymerization of vinyl acetate to poly(vinyl acetate) (PVAc), and subsequent hydrolysis of PVAc gives PVA [12].
PVA can be crosslinked using several methods, such as the use of crosslinking chemical agents, using an electron beam or γ-irradiation, or the physical crosslinking due to crystallite formation. For biomedical applications, physical crosslinking has the advantages of not leaving residual amounts of the toxic crosslinking agent, and higher mechanical strength than the PVA gels crosslinked by either chemical or irradiative techniques [22, 23]. In chemical cross-linking, the chemical agents that react with the hydroxyl groups are glutaraldehyde, ethylaldehyde, terephthalaldehyde, formaldehyde, hydrochloric, boric or maleic acid, among others [11, 24]. Physical crosslinking forms a hydrogel with a network of semi-crystallites of hydrogen bonds of polymer filled with solvent [25]. It has been shown that the mechanical properties of the hydrogels, including elasticity and strength, can be altered by changing the PVA concentration, the number of freeze/thaw cycles, the process thawing rate, the freezing holding time, and the freezing temperature [11, 26, 27]. Increasing the PVA concentration results in hydrogels with higher crystallinity and added stability upon swelling, which increases its tensile strength and tear resistance. The lower the initial concentration of PVA, the fewer the polymer chains in solution, and there may be less number of crystalline regions created in the cycled PVA. Increasing the number of freeze/thaw cycles increases the strength and stiffness of the hydrogel by reinforcing existing crystals within the structure [11, 28, 13]. Decreasing the thawing rate of frozen PVA solutions increases the tensile strength because the solutions are kept for longer periods at temperatures below 0° C., allowing for increasing movements of polymer chains which result in further entanglements and increased crystallite size and numbers. The freezing holding time also has a drastic effect, with samples frozen up to 10 days giving the most mechanically strong PVA hydrogels [24, 13, 25, 27]. The freezing temperature has an interesting effect. The freezing temperature controls the phase equilibria and dynamics, where the lower the temperature of the system the lower the amount of unfrozen solvent in the liquid regions. Therefore, the lower the temperature the less opportunity for chain mobility in the polymer rich regions, giving less chances of crystallite growth and formation. This explains why keeping the frozen PVA solutions at −10° C. produces somewhat more rigid hydrogels than those kept for the same period of time at −20 or −30° C. The freezing rate was shown not to have drastic effects on the properties of the hydrogel [11, 13, 25]. PVA hydrogels not only have tensile strength and elongation, but also flexibility and elasticity. Research has proven its ability to recover to its original shape after being deformed to strains of 50%, showing excellent persistence and repeatability of the recovery [25].
Physical crosslinking allows the PVA hydrogels to retain their original shape and be extended up to six times their size. This behaviour shows its rubbery and elastic nature and the high mechanical strength [29, 26]. There are various theories proposed in the literature to explain why thermal cycling increases the elastic modulus of PVA. The most accepted theory describes the physical cross-linking process as an entropic reordering phenomena. Water is likely to bind to the polymer by hydrogen bonding. When the solution freezes, ice crystals force the polymer chains close to each other forming high local polymer concentration regions or nuclei. When the material thaws, these nuclei act as crosslinking sites for polymers molecules, which realign and form hydrogen bonds to form crystallites and polymer chain entanglements. The crystalline regions are formed within the polymer rich regions, with further cycling increasing both the size and number of the crystalline regions by repeating the process [11, 30, 29]. On a molecular level, the crystallites of PVA can be described as layered structure, with a double layer of molecules held together by hydroxyl bonds, while weaker van der Waals forces-operate between the double layers. This folded chain structure leads to ordered regions (crystallites) within an unordered, amorphous polymer matrix [12]. The mechanical properties of PVA are very unique compared to other polymers. The stress-strain curves for the polymeric materials are initially linear and then curve towards the strain axis. On the other hand, the PVA curve displays an exponential stress-strain curve similar to the characteristics of soft biological tissues, with the curve shifting towards the stress axis.
PVA materials have been reported to be ideal candidates as biomaterials, due to their high degree of swelling, uncomplicated chemical structure, rubbery/elastic nature, non-toxic, non-carcinogenic, and bioadhesive characteristics. Some of the biomedical applications include tissue reconstruction and replacements, cell entrapment and drug delivery, soft contact lens material, wound covering bandage for burn victims, quality control phantom for MR, among other medical applications [30, 12].
Although PVA hydrogel can be processed to possess mechanical properties similar to some soft biological tissues, there are tissues such as heart valve cusps and cartilage that have mechanical properties that are beyond the range of the low temperature processed PVA. Also, for medical device applications, for durability, the most ideal material would be one that has mechanical properties that mimic the soft tissue to be replaced within the physiological range but stronger beyond this range. These requirements imply that a material more than PVA is required for good, durable medical device applications. One approach is to create a PVA based composite that possesses the properties requirements outlined.
Therefore there is a need for a composite material that has properties similar to that of natural tissue for medical device applications. Moreover, if would be a further advantage if such material is capable of delivering bioactive agent locally where the device is implanted.
Bacterial Cellulose
Bacterial cellulose has many characteristics that make it valuable for biomedical applications, including its polyfunctionality, hydrophilicity, and biocompatibility [33]. Cellulose is a linear polymer made of glucose molecules linked by β(1-4) glycosidic linkages. Its chemical formula is (C6H10O5)n. There are four principle sources of cellulose. The majority of cellulose is isolated from plants. A second source is the biosynthesis of cellulose by different microorganisms, including bacteria (acetobacter, aerobacter, pseudomonas), algae, and fungi among others. The other two less common sources include the enzymatic in vitro synthesis starting from cellobiosyl fluoride, and the chemosynthesis from glucose by ring-opening polymerization of benzylated and pivaloylated derivatives [31, 32]. Cellulose is not uniformly crystalline, but ordered regions are extensively distributed throughout the material, and these regions are called crystallites. The long cellulose chains lie side by side held together by hydrogen bonds between the hydroxyl groups. These chains are twisted into structures called microfibrils, which are twisted into fibers [33, 31].
Bacterial cellulose is produced by strains of the bacterium Acetobacter xylinum, which is typically found on decaying fruits, vegetables, vinegar, fruit juices, and alcoholic beverages. It is a Gram-negative, rod shaped and strictly aerobic bacterium. Bacterial cellulose produced has very high purity and contains no lignin, hemicelluloses, pectin, and waxes as plant cellulose does. Therefore, production of bacterial cellulose has the advantage of not requiring the harsh chemical treatment needed for plant cellulose production. This chemical treatment also has the disadvantage of altering the natural structural characteristics of cellulose [33, 31, 32]. Bacterial cellulose differs from plant cellulose with respect to its high crystallinity, ultra-fine network structure, high water absorption capacity, high mechanical strength in the wet state, and availability in an initial wet state [32]. Bacterial cellulose pellicles are formed in static culture. The pellicle has an ultra-fine network structure of ribbons 500 nm wide and 10 nm thick. The ribbons consisted of smaller microfibrils with a width of around 3 nm and a fiber diameter of less than 130 nm compared to the over 14 mm found in birch [31, 32]. Bacterial cellulose including the pellicle possesses a high water retention capacity. Water retention values can reach up to 1000%, which are significantly higher than that for plant cellulose. The water retention is drastically decreased after air-drying the bacterial cellulose and reswelling in water, with values comparable to those of plant cellulose [31, 32].
Bacterial cellulose can also be prepared in shake culture in flasks and in agitated culture in a bioreactor. These approaches are more efficient methods for bacterial cellulose production and are preferred for large scale production of bacterial cellulose.
Bacterial cellulose, being a hydrophilic, highly water swollen and biocompatible natural polymer which is ideally suited to be the reinforcing fibers in the preparation of a composite material for soft tissue replacement devices. Such composite material can be created when it is used in combination with PVA.
Fiber reinforced composites provide improved strength, stiffness, and fatigue resistance. The softer, more elastic matrix transmits the force to the fibers, which normally carry most of the applied force. The modulus of elasticity and strength of the composite depend on various factors. The fibers can be short, long, or continuous with typical diameters in the range of 10 to 150 microns. The larger the aspect ratio (length/diameter) of the fibers, the higher the strength of the composite. The greater the fiber volume fraction also increases the composite strength and stiffness up to 80%. The orientation of the fibers is also an important factor. Short, randomly orientated fibers give relatively isotropic behaviour. Long, unidirectional arrangements of fibers produce anisotropic properties, with good strength and stiffness in the orientation parallel to the fibers. The raw fiber properties are important, with strong, stiff, and lightweight fibers being the most commonly used. The matrix properties are also important, supporting the fibers, keeping them in the proper position, transferring the load to the fibers, and preventing cracks in the fibers. Therefore, good bonding between the fibers and the matrix is required for the successful transfer of load in the composite [34]. Joining two or more materials may give composites with properties not attainable by the original materials. The materials are selected to improve properties such as stiffness, strength, corrosion resistance, high-temperature performance, and conductivity [34].
Uryu [35] reported the formation of a biodegradable polymeric material that can be decomposed in soil. The bacterial cellulose (with ribbon shaped micro-fibrils) that can be biologically decomposed by microbes was mixed with a biodegradable polymeric material to produce an improved composite with higher tensile strength. The bacterial cellulose was produced in a liquid culture medium using different types of microbes, including Acetobacter xylinum, collected and dried into a powdery state and mixed with the polymer to produce the composite. Various polymers were used, including PVA. The composites ranged from bacterial cellulose concentrations as low as 1% to 99%. The final composite was dried and used for high-strength cabinets for audio/video apparatus. After the lifetime of the device is reached, the composite material can be buried in the ground for waste disposal and it is eventually decomposed to protect the environment.